Method and fuel cell for electrochemical measurement of analyte concentration in vivo

ABSTRACT

The invention relates to a method for the electrochemical measurement of an analyte concentration in vivo, comprising a fuel cell with which the analyte to be measured is reacted catalytically with an enzyme contained in an enzyme layer and which supplies an electrical voltage, dependent on the analyte concentration to be measured, beween an anode and a cathode, which voltage is measured. In the catalytic reaction of the analyte to be measured in the enzyme layer, a product is generated which, as fuel of the fuel cell, oxidizes on the anode and is reduced on the cathode. The invention further relates to a fuel cell for such a method.

RELATED APPLICATIONS

This application is a continuation of PCT/EP2010/006770, filed Nov. 6,2010, which claims priority to EP10002588.1, filed Mar. 11, 2010, bothof which are incorporated herein by reference in their entirety.

BACKGROUND

The invention relates to a method for the electrochemical measurement ofan analyte concentration in vivo by means of a fuel cell whichcatalytically converts the analyte to be measured with an enzymecontained in an enzyme layer and supplies an electrical voltage betweenan anode and a cathode.

As is known from U.S. Pat. No. 3,837,339, fuel cells are used aselectrochemical sensors for in vivo measurement of analyteconcentrations, for example, glucose concentration. Therein, the analyteto be measured is catalytically converted as the fuel of the fuel cell.The fuel cell described in U.S. Pat. No. 3,837,339 oxidizes glucose onthe anode, thereby generating gluconic acid. Oxygen is reduced on thecathode.

In order to improve the efficiency of these reactions, it is known toprovide the cathode with an enzyme layer which contains an enzyme forcatalytically converting the analyte and to apply an enzyme layercontaining an enzyme to the anode for the catalytic reduction of oxygen.For example, such fuel cells are known from U.S. Publication No.2005/0118494. In these fuel cells, the anode is covered with an enzymelayer which contains glucose oxidase and the cathode is covered with anenzyme layer which contains laccase. In this manner, the anode reaction,i.e. the oxidation of glucose, and the cathode reaction, i.e. thereduction of oxygen, can be accelerated.

A problem arising in connection with the in vivo measurement of analyteconcentrations with such sensors is that the intensity of the measuringsignal can be affected by an oxygen deficiency in the environment of thecathode. In particular in the event of prolonged operation, depletion ofthe oxygen concentration can occur in the environment of the cathode,the oxygen concentration being already subject to great variations insubcutaneous fatty tissue. This may falsify the measuring signal. Afurther problem is that intermediate products that are harmful to healthor even toxic, for example, hydrogen peroxide, can often develop in thecatalytic conversion of analytes. If such intermediate products exitfrom the sensor, this may result in inflammation and require prematureremoval of the sensor.

SUMMARY

The present invention provides a method of measuring analyteconcentrations over a prolonged period of time in a cost-effective andprecise manner.

In exemplary embodiments, a product is generated in the enzyme layer bycatalytic conversion of the analyte to be measured, said product thenoxidizing on the anode and being reduced on the cathode. In this manner,a measuring signal independent of the local oxygen concentration can beachieved because the oxygen reduction reaction on the cathode, whichoccurs in prior art devices, is no longer required. For this reason, themethod according to these teachings allows measuring analyteconcentrations in vivo with improved precision over a prolonged periodof time.

With the method taught herein, it is, for example, possible to generatehydrogen peroxide as a product by catalytic conversion of the analyte tobe measured, wherein said hydrogen peroxide subsequently oxidizes on theanode and can be reduced on the cathode. By converting the hydrogenperoxide, which is problematic in terms of health, both on the anode andon the cathode, the concentration thereof is, typically, only half ashigh as in conventional sensors which convert hydrogen peroxide only atone electrode. For this reason, harmful effects of the hydrogen peroxidecan be avoided to a far better degree than is the case with conventionalsensors. As an alternative to hydrogen peroxide, use can also be made ofother redox amphoteric substances, such as sulfites and aldehydes. Anexample of a redox amphoteric aldehyde is 5-hydroxy indolyl acetaldehydewhich can be generated from serotonin using monoamine oxidases and whichoxidizes to 5-hydroxy indolyl acetic acid and can also be reduced to5-hydroxytryptophol.

For the method according to this disclosure, an electrochemical sensorwhich comprises an anode and a cathode is used for in vivo measurementof an analyte concentration. The surface of the cathode is formed of adifferent material than that of the anode. The anode, the cathode and anenzyme layer which contains an enzyme for the catalytic conversion ofthe analyte to be measured form a fuel cell. The enzyme layer of thesensor forms a diffusion path for a product generated by catalyticconversion of the analyte, with the result that the sensor oxidizes thisproduct on the anode and reduces it on the cathode.

A reaction competing with the reduction of the product generated bycatalytic conversion of the analyte may be the reduction of oxygen.Practically, however, this competitive reaction can be disregardedbecause the reduction of the product generated by catalytic conversionof the analyte predominates. If the product is hydrogen peroxide, thereaction rate of the reduction of oxygen, typically, is less than atenth, normally even less than a hundredth of the reaction rate of thereduction of hydrogen peroxide. In other words, more than 10, normallyeven more than 100 molecules of hydrogen peroxide are reduced for onemolecule of oxygen. In a sensor according to these teachings, theproduct is, therefore, oxidized on the anode and reduced on the cathodeas a signal determining reaction. This means that the signal contents ofany competitive reactions are less than 10 percent of the signalintensity.

On the one hand, the fact that the oxygen reduction can be disregardedis based on an overpotential of oxygen. Such an overpotential alwaysoccurs in redox systems to a greater or lesser extent. The reductionreaction of oxygen is inhibited by this overpotential and, therefore,takes place at a rate that is significantly slower than the reduction ofthe product generated by catalytic conversion of the analyte. On theother hand, the standard electrode potential of oxygen, which is only1.22 V, is relatively small, in particular in comparison with thestandard electrode potential of hydrogen peroxide, which is 1.77 V. Forthis reason, the reduction of oxygen is thermodynamicallydisadvantageous in this respect as well.

A fuel cell according to this disclosure may not only be used as asensor for the measurement of an analyte concentration but also as anenergy source of an implanted device in the body of a patient, forexample, for a cardiac pacemaker.

In a fuel cell according to this disclosure, the enzyme layer forms adiffusion path for the product generated by catalytic conversion of thesubstance or the analyte, with the result that this product oxidizes onthe anode and is reduced on the cathode while the sensor is in use. Thiscan be most easily achieved by the enzyme layer touching both the anodeand the cathode. For example, the anode and the cathode can be arrangedon a common substrate, wherein the enzyme layer covers both the anodeand the cathode. A further possibility is to arrange the enzyme layerbetween the anode and the cathode. In particular in case of a sandwichedarrangement, further intermediate layers may be provided between theenzyme layer and the anode or between the enzyme layer and the cathode,provided these intermediate layers are permeable to the redox amphotericproduct, for example, hydrogen peroxide.

Inasmuch as a sensor is mentioned in the following description, thefeatures referred to therein can also be used in a fuel cell accordingto these teachings, which is used as an energy source of a cardiacpacemaker or another device in the body of a patient. In particular, asensor as taught herein can be readily used as an energy source for acardiac pacemaker.

The exemplary sensor can be produced in a very simple and cost effectivemanner because, in essence, the only items required are an anode, acathode and the enzyme layer. The expenditure connected with theproduction of separate enzyme layers for anode and cathode can beavoided.

The enzyme molecules that are typically included in the enzyme layer arean oxidase, for example glucose oxidase or lactose oxidase, but can, forexample, also be a hydrogenase. The enzyme molecules can be covalentlybonded in the enzyme layer. It is, however, also possible that theenzyme molecules are only admixed to the material of the enzyme layerand are mobile in the enzyme layer. Particularly in this case, theenzyme layer can be covered with a top layer, for example, with apolymer film, which is impermeable to enzyme molecules. Suitable toplayers which are permeable to water and the analyte molecules to bemeasured can, for example, be produced from polyurethanes, polyvinylchlorides, polycarbonates, polytetrafluoroethylenes, acrylates orsilicones. Sulfonated tetrafluoroethylenes, such as they arecommercially available under the brand name of Nafion, are particularlysuitable.

For the enzyme layer, use can, in particular, be made of polymers thatare permeable to water. Particularly suitable are plastics, for example,polyurethane. Further possibilities are, for example, pectin, gelatin orother natural materials that are permeable to water. The enzyme layercan be formed as a plastic matrix with included enzyme molecules.

Ideally, the enzyme layer consists of a material that is an insulator inits dry state. For operation of the sensor, the enzyme layer takes upwater and becomes a ionic conductor. It is desirable that theintermediate product has a high mobility inside the enzyme layer.

The cathode of the sensor according to these teachings can have ametallic surface, for example of palladium. However, other noble metalsand noble metal alloys are also suitable. In this manner, the cathodecan be cost effectively formed as a conducting track on a substrate ofplastic. The cathode can also be designed with a non-metallic surfacewhich allows a reduction of the intermediate product.

The anode can, likewise, have a metallic surface provided it consists ofa different metal than that of the cathode surface. However, the anodetypically has a non-metallic surface, for example a surface thatcontains carbon particles or graphite particles. The non-metallicsurface of the anode can be arranged as a covering on a metallicconductor.

Carbon black or graphite particles can be easily mixed with a polymericbinding agent to form a paste into which catalytically active particlescan be blended. Suitable blendings allow the generation of pastes ofvarying compositions, said pastes then serving to form the cathodeand/or anode on metallic conducting tracks.

Preferably, a conductor running to the anode is connected to a conductorrunning to the cathode via an electrical resistor. This resistor causesa continuous electric current between the anode and the cathode, withthe result that the redox amphoteric product is converted continuously,this being an important advantage in particular in case of products thatare problematic in terms of health, such as hydrogen peroxide.

A further advantage is that, in contrast to conventional amperometrichydrogen peroxide sensors, no external voltage supply is required inorder to continuously convert hydrogen peroxide.

The electrical resistor can, for example, be formed as a conductingtrack on a substrate on which there is also the conductor running to theanode and the conductor running to the cathode. The conducting trackused for forming the resistor can be made of the same material as thatof the conducting tracks running to the anode and the cathode, whereintheir diameter, in particular their thickness and/or their width, isselected appropriately small in order to obtain the desired value of theresistor. However, the resistor can be made of a different material thanthat of the conductors running to the anode and the cathode. Forexample, a conducting track can be made as a paste of a resistormaterial and printed onto the substrate. The resistor also can be amaterial mixture which contains carbon particles, for example carbonblack and/or graphite, and a binding agent. For example, a paste thatcontains carbon particles and solidifies after having been applied canbe a suitable resistor material. The resistor typically has a value ofat least 1 megohm. As a general rule, however, resistance values inexcess of 1 gigaohm are not advantageous. As a general rule, resistancevalues between 10 megohms and 100 megohms are advantageous.

The electrical resistor can, however, also be integrated into a plugcontact which is connected to both conducting track ends in a conductingmanner during operation of the sensor.

An advantageous refinement provides that both the reduction and theoxidation of the redox amphoteric product which is generated byenzymatic conversion of the analyte to be measured are supported byelectrochemically active catalysts. Catalysts that are suitable for theoxidation of hydrogen peroxide are, for example, metal oxides, inparticular manganese dioxide, and metallo-organic compounds, for examplecobalt phthalocyanine. In particular, manganese dioxide in powder formcan be easily mixed with graphite particles and, in this manner, beintegrated into the surface of the anode.

BRIEF DESCRIPTION OF THE DRAWINGS

Further details and advantages of this disclosure will be illustrated bymeans of embodiments with reference being made to the accompanyingdrawings. Therein, identical parts or parts that are corresponding toeach other are designated with consistent reference symbols.

FIG. 1 shows an embodiment of a sensor according to the invention;

FIG. 2 is a sectional view of FIG. 1;

FIG. 3 shows an example of the graph of the voltage measured with thesensor in mV (left-hand axis) and of the hydrogen peroxide concentrationin mmol (right-hand axis) over time;

FIG. 4 shows a measurement example of the voltage delivered by thesensor in mV in relation to the glucose concentration in mmol;

FIG. 5 shows a further embodiment of a sensor according to theinvention;

FIG. 6 is a sectional view of the sensor shown in FIG. 5;

FIG. 7 shows a further embodiment of a sensor according to theinvention;

FIG. 8 is a longitudinal sectional view of a further embodiment of asensor according to the invention;

FIG. 9 is a longitudinal sectional view of a further embodiment of asensor according to the invention;

FIG. 10 is a sectional view taken from cutting line AA of FIG. 9;

FIG. 11 is a sectional view taken from cutting line BB of FIG. 9;

FIG. 12 is a longitudinal sectional view of a further embodiment of asensor according to the invention;

FIG. 13 is a sectional view taken from cutting line AA of FIG. 12;

FIG. 14 is a sectional view taken from cutting line BB of FIG. 12;

FIG. 15 is a sectional view taken from cutting line CC of FIG. 12;

FIG. 16 shows a further exemplary embodiment of a sensor according tothe invention;

FIG. 17 is a sectional view of FIG. 16; and

FIG. 18 shows a further embodiment of a sensor according to theinvention.

DETAILED DESCRIPTION

The embodiments described below are not intended to be exhaustive or tolimit the invention to the precise forms disclosed in the followingdetailed description. Rather, the embodiments are chosen and describedso that others skilled in the art may appreciate and understand theprinciples and practices of this disclosure.

The sensor shown in FIG. 1 works according to the principle of a fuelcell. A redox amphoteric product which forms the fuel of the fuel cellis generated from the analyte to be measured, for example glucose orlactate. For this reason, the energy supplied by the fuel cell becomeshigher as the analyte concentration to be measured increases. Theelectrical voltage drop across a load resistor between the anode and thecathode can, therefore, be used as a measuring signal for determiningthe analyte concentration.

The sensor shown in a top view in FIG. 1 and in a sectional view in FIG.2 has an anode 1 and a cathode 2 which are covered by a common enzymelayer 3. The anode 1 and the cathode 2 are each disposed at the end of aconducting track 5, 6 arranged on a substrate 4, for example a plasticsheet. The conducting tracks 5, 6 can consist of a noble metal, forexample palladium, which may also form the surface of the cathode 2. Theanode 1 can be formed as a covering of the conducting track 5, forexample made of carbon particles and a binding agent. The two conductingtracks 5, 6 are connected to an electrical load resistor 7 and arecovered by an electrically insulating layer 8 that is impermeable towater. The ends of the conducting tracks 5, 6 project from under theinsulating layer 8. Hence, the contact surfaces 5 a, 6 a are not coveredby the electrically insulating layer 8 that is impermeable to water,just as is the case with the anode 1 and the cathode 2.

The analyte molecules, for example glucose molecules, that are diffusinginto the enzyme layer 3 that is permeable to water are enzymaticallyconverted by the enzyme molecules contained in the enzyme layer 3, forexample an oxidase, whereby a redox amphoteric product, for examplehydrogen peroxide, is generated. The redox amphoteric product is mobilein the enzyme layer 3 and, therefore, arrives both at the anode 1 andthe cathode 2. The redox amphoteric product is oxidized on the anode 1and reduced on the cathode 2. In order to promote the oxidationreaction, a catalyst, for example manganese dioxide, can be admixed tothe anode material. Since the redox amphoteric product iselectrochemically converted on the anode 1 and on the cathode 2, anelectrical voltage develops between the anode 1 and the cathode 2. Theelectrical voltage between the anode 1 and the cathode 2 is measuredacross the load resistor 7 and used as a measuring signal fordetermining the analyte concentration. The end 5 a and 6 a facing awayfrom the anode 1 and the cathode 2 can be broadened in order to act ascontact pads facilitating the connection of a voltage meter.

In an ideal fuel cell, the voltage between the anode 1 and the cathode 2is only dependent on the electrochemical potentials which develop as aresult of the anode and cathode reactions and on the size of the loadresistor 7 between the anode 1 and the cathode 2. For this reason, theelectrical voltage between the anode 1 and the cathode 2 across the loadresistor 7 develops as a result of the reaction rates of the reaction onthe anode 1 and the reaction on the cathode 2.

Since, in the sensor shown, the same substance, e.g., hydrogen peroxide,is converted at both the anode 1 and the cathode 2, both the anodereaction and the cathode reaction are, essentially, determined by theanalyte concentration in the enzyme layer 3. That is to say, the rate atwhich the product is generated, which is converted at the anode 1 andthe cathode 2, is approximately proportional to the analyteconcentration within a wide concentration range.

The load resistor 7 can be formed by a conducting track of resistormaterial, for example a paste containing graphite particles, whichconnects the conducting track 5 running to the anode 1 to the conductingtrack 6 running to the cathode 2. Preferably, the resistor 7 is arrangedbelow the insulating layer 8. In principle, however, it is also possibleto use the resistor 7 to connect the connection-sided ends of theconducting tracks 5, 6 projecting from under the insulating layer 8.

FIG. 3 shows a measurement example of the voltage supplied by the sensorin millivolts along with the hydrogen peroxide concentration inmillimoles over the time t in seconds. Therein, the left-hand ordinateindicates the voltage U in millivolts for the measurement curve A, andthe right-hand ordinate indicates the hydrogen peroxide concentration inmillimoles for the associated concentration graph which is representedby curve B. As can be seen, a new equilibrium voltage develops betweenthe anode 1 and the cathode 2 across the load resistor 7 within a fewseconds when the hydrogen peroxide concentration rises in a step-likemanner.

FIG. 4 shows a measurement example of the voltage U supplied by thesensor in millivolts in relation to the glucose concentration inmillimoles/liter. Therein, it can be seen that the electrical voltagebetween the anode and the cathode is higher as glucose concentrationincreases. For this reason, the associated glucose concentration can bedetermined with a calibration curve based on the measured voltage. Ananalyte concentration can, therefore, be determined by measuring thevoltage dropping at a resistor 7 which connects a conductor 5 running tothe anode to a conductor 6 running to the cathode.

FIG. 5 shows a further embodiment of a sensor according to theinvention. FIG. 6 is a longitudinal sectional view of FIG. 5. Inessence, the sensor shown in FIGS. 5 and 6 differs from the embodimentshown in FIGS. 1 and 2 only by a covering layer 9 covering the enzymelayer 3. The covering layer 9 is permeable to analyte molecules as wellas water and can fulfill a plurality of functions which are each leadingto an improvement of the sensor but are not necessarily required.

For example, the covering layer 9 can be impermeable to enzymemolecules. By counteracting an exit of enzyme molecules from the sensor,the compatibility of the sensor can be improved because exiting enzymemolecules might have harmful effects in the body tissue of a patient.This function of the covering layer 9 is, in particular, to advantagewhenever the enzyme molecules in the enzyme layer 3 are not covalentlybonded. The enzyme layer 3 can contain enzyme molecules that arecovalently bonded. For example, enzyme molecules can be covalentlybonded to polymers of a matrix. It is, however, also possible that theenzyme molecules are only admixed to the material of the enzyme layer 3and can diffuse therein. Particularly in the latter case, a coveringlayer 9 that is impermeable to enzyme molecules is a significantadvantage. For this purpose, the covering layer 9 can, for example, beproduced from polyurethanes, polyvinyl chloride, polycarbonate,polytetrafluoroethylene, polyacrylates, silicones, polyvinyl pyrrole ormixtures of such polymers.

Advantageously, the covering layer 9 can also form a reservoir foranalyte molecules. Analyte molecules can flow from said reservoir to theenzyme layer 3 in the event of a temporary failure of the fluid exchangein the environment of the sensor. If the exchange of body fluid is,temporarily, restricted or even prevented in the environment of thesensor, for example caused by movements of the patient's body, analytemolecules stored in the covering layer 9 can continue to diffuse to theenzyme layer 3. In this manner, the covering layer 9 can have the effectthat a noticeable depletion of the analyte concentration and acorresponding falsification of the measurement results will not occurbefore a considerably longer time interval has elapsed. For thispurpose, the covering layer 9 can also have a considerably greaterthickness than is suggested in FIG. 6, which is not true to scale.

A further function of the covering layer 9 can be to provide a diffusionresistance for the analyte to be measured, i.e., act as a diffusionbarrier. Due to its diffusion resistance, the covering layer 9 has theeffect that a lesser number of analyte molecules arrive at the enzymelayer 3 per time unit. By means of the covering layer 9, the rate atwhich analyte molecules are converted can, therefore, be reduced and adepletion of an analyte concentration in the environment of the sensor,thus, be counteracted.

A further embodiment of a sensor according to the invention is shown inFIG. 7. FIG. 8 is a longitudinal sectional view of FIG. 7. In essence,this embodiment differs from the embodiment of FIGS. 1 and 2 only inthat the enzyme layer 3 does not cover the anode 1 and the cathode 2.Instead, the enzyme layer 3 is arranged between the anode 1 and thecathode 2. In this arrangement as well, the enzyme layer 3 forms adiffusion path for the analyte and the redox amphoteric product formedby the conversion thereof, for example hydrogen peroxide, to the anode 1and to the cathode 2. As is the case in the remaining embodiments, theproduct is, therefore, oxidized on the anode 1 and reduced on thecathode 2.

The embodiment shown in FIGS. 7 and 8 can be provided with a coveringlayer 9, as has been illustrated above with respect to the embodimentshown in FIGS. 5 and 6. In the embodiments of FIGS. 7 and 8, a possiblecovering layer 9, preferably, does not only cover the enzyme layer 3 butalso the anode 1 and the cathode 2.

FIG. 9 is a longitudinal sectional view of a further embodiment of asensor according to the invention. FIG. 10 is a sectional view takenfrom cutting line AA of FIG. 9. FIG. 11 is a sectional view taken fromcutting line BB of FIG. 9.

In this embodiment, the anode 1 and the cathode 2 are arranged ondifferent substrates 4 a, 4 b. The substrates 4 a, 4 b can, for example,be formed as plastic sheets and each support a conducting track 5, 6,with the anode 1 and the cathode 2, respectively, being disposed at theends of said conducting track 5, 6. The enzyme layer 3 is arrangedbetween the two substrates 4 a, 4 b. The front face of this sandwichedarrangement is covered by a covering layer 9, such as it was illustratedin the context of the embodiment of FIGS. 5 and 6. As is the case withthe embodiments described above, the conducting tracks 5, 6 can,likewise, be covered by an insulating layer 8. In addition, a spacer 10can be arranged between the two substrates 4 a, 4 b.

In this embodiment, the load resistor 7 that connects the two conductingtracks 5, 6 can be a separate component that is arranged between thesubstrates 4 a, 4 b. For example, this component can be soldered to theconducting tracks 5, 6 or be connected thereto in a clamping manner. Inparticular, it is also possible that the spacer 10 supports the loadresistor 7. For example, the resistor element 7 can be applied as aconducting paste into the intermediate space between the two substrates4 a, 4 b and onto the front face of the spacer 10, with the result thatthe load resistor 7 comes into contact with the two conducting tracks 5,6.

FIG. 12 is a longitudinal sectional view of a further exemplaryembodiment of a sensor. FIG. 13 is a cross-sectional view of this sensortaken from cutting line AA plotted in FIG. 12. FIG. 14 is across-sectional view of this sensor taken from cutting line BB plottedin FIG. 12. FIG. 15 is a cross-sectional view of this sensor taken fromcutting line CC plotted in FIG. 12.

In this embodiment, the anode 1 is arranged on the inner side of asleeve and the cathode 2 is arranged on a conductor surrounded by thesleeve. This embodiment can be modified to the effect that the cathode 2is arranged on the inner side of the sleeve and the anode is arranged onthe conductor surrounded by the sleeve. For example, the conductorsurrounded by the sleeve can be a wire. The sleeve can be made of metalor be a small plastic tube the inner side of which has a metalliccoating in order to form an electrical conductor running to theelectrode arranged on its inner side, preferably to the anode 1.

Similar to the embodiments described above, the conductors 5, 6 runningto the anode 1 and the cathode 2 are covered with an insulating layer 8.The enzyme layer 3 is disposed between the anode 1 and the cathode 2. Inthe embodiment shown, the remaining interior region of the sensor is, inessence, completely filled with the enzyme layer 3, for example with aplastic matrix containing enzyme molecules. The front face of the sensoris covered with a covering layer 9, such as it can also be present inthe embodiment illustrated above.

In the embodiment shown, the enzyme layer 3 extends to the load resistor7 which connects the anode conductor to the cathode conductor. However,it can be to advantage to fill a section of the sleeve with insulatingmaterial. In this case, the load resistor 7 can be arranged spaced apartfrom the enzyme layer 3, for example on the insulating material closingthe sleeve. In this embodiment as well, the load resistor 7 can be,advantageously, formed as a paste which contains carbon particles andsolidifies after having been applied. It is, however, also possible toform the load resistor 7 as a conventional resistor element which is,for example, soldered to the anode conductor and the cathode conductoror is conductively connected thereto in any other manner.

FIG. 16 shows a further embodiment of an exemplary sensor. FIG. 17 is asectional view of FIG. 16. In essence, this embodiment differs from theembodiment shown in FIGS. 5 and 6 only in that two anode-cathode pairs 1a, 2 a and 1 b, 2 b are disposed therein. The conductor 6 is running tothe anode 1 a, the conductor 5 is running to the cathode 2 a, theconductor 6 b is running to the anode 1 b, and the conductor 5 b isrunning to the cathode 2 b.

For this reason, the sensor shown has two fuel cells, with the resultthat, in principle, there are two different sensors. The two fuel cellscan be provided for the measurement of the same analyte or, by usingdifferent enzymes, can be intended for the measurement of differentanalytes. For example, one of the two fuel cells can be provided for themeasurement of glucose and the other fuel cell can be provided for themeasurement of lactates. If both fuel cells are intended for themeasurement of the same analyte, it is to particular advantage if thetwo fuel cells each form a sensor with different sensitivity. Bymeasuring low analyte concentrations with one of the fuel cells and highanalyte concentrations with the other fuel cell, it is possible toachieve a measurement precision that is higher as a whole. For example,different measurement sensitivities can be realized by different enzymeconcentrations in the enzyme layers. A further possibility is to coverthe enzyme layers with top layers 9 that are permeable to differentdegrees.

A further embodiment of a sensor is shown in FIG. 18. Similar to theembodiment shown in FIGS. 16 and 17, there are two different fuel cellsin the sensor shown in FIG. 18 as well. In this embodiment, however, thetwo fuel cells are not arranged on the same side of a substrate.Instead, one of the two fuel cells is arranged on the top side and theother fuel cell on the bottom side of the substrate 4.

While exemplary embodiments incorporating the principles of the presentinvention have been disclosed hereinabove, the present invention is notlimited to the disclosed embodiments. Instead, this application isintended to cover any variations, uses, or adaptations of the inventionusing its general principles. Further, this application is intended tocover such departures from the present disclosure as come within knownor customary practice in the art to which this invention pertains andwhich fall within the limits of the appended claims.

REFERENCE SYMBOLS

-   1, 1 a, 1 b Anode-   2, 2 a, 2 b Cathode-   3 Enzyme layer-   4 Substrate-   5, 5 b Conductor-   5 a Contact surface-   6, 6 b Conductor-   6 a Contact surface-   7 Load resistor-   8 Insulating layer-   9 Covering layer-   10 Spacer

What is claimed is:
 1. A method for electrochemical measurement of ananalyte concentration in vivo, comprising: providing a fuel cellcomprising an enzyme layer, an anode and a cathode; providing to thefuel cell a sample fluid having the analyte; using the fuel cell tocatalytically convert the analyte to be measured with an enzymecontained in the enzyme layer, wherein the catalytic conversiongenerates a fuel for the fuel cell which is oxidized on the anode and isreduced on the cathode; and generating an electrical voltage between theanode and the cathode, said voltage corresponding to the concentrationof the analyte.
 2. The method of claim 1, wherein the oxidation of thefuel is supported by a catalyst.
 3. The method of claim 2, wherein thecatalyst comprises a metal oxide or a metallo-organic compound.
 4. Themethod of claim 1, wherein the anode and the cathode are connected toeach other via an electrical resistor and the electrical voltage betweenthe anode and the cathode is measured as a voltage drop at the resistor.5. A fuel cell, comprising: an anode and a cathode, the cathode having asurface made of a different material than that of the anode; an enzymelayer containing an enzyme for the catalytic conversion of a substanceoccurring in the human body, the enzyme layer forming a diffusion pathto the anode and to the cathode for a fuel product generated bycatalytic conversion of the substance, the fuel cell adapted to oxidizethe fuel product on the anode and reduce it on the cathode.
 6. Anelectrochemical sensor for in vivo measurement of an analyteconcentration comprising a fuel cell according to claim 5, wherein thesubstance for the catalytic conversion is the analyte to be measured. 7.The electrochemical sensor of claim 6, wherein the enzyme layer contactsthe anode and the cathode.
 8. The electrochemical sensor of claim 6,wherein the anode and the cathode are arranged on a common substrate andthe enzyme layer covers both the anode and the cathode.
 9. Theelectrochemical sensor of claim 6, wherein the cathode has a metallicsurface.
 10. The electrochemical sensor of claim 6, wherein the anodehas a non-metallic surface.
 11. The electrochemical sensor of claim 6,wherein the enzyme is an oxidase.
 12. The electrochemical sensor ofclaim 6, wherein the fuel is hydrogen peroxide.
 13. The electrochemicalsensor of claim 6, wherein the enzyme layer is an insulator in its drystate.
 14. The electrochemical sensor of claim 6, wherein an electricalresistor connects a first conductor running to the anode to a secondconductor running to the cathode.
 15. The electrochemical sensor ofclaim 14, wherein the resistor is a mixture which contains carbonparticles and a binding agent.
 16. The electrochemical sensor of claim6, wherein the enzyme layer is covered by a covering layer.